Method and device for monitoring and improving patient-ventilator interaction

ABSTRACT

Method and apparatus for non-invasively determining the time onset (T onset ) and end (T end ) of patient inspiratory efforts. A composite pressure signal is generated comprising the sum of an airway pressure signal, a gas flow pressure signal obtained by applying a gain factor (K f ) to a signal representing gas flow rate and a gas volume pressure signal obtained by applying a gain factor (K v ) to a signal representing volume of gas flow. K f  and K v  values are adjusted to result in a desired linear trajectory of composite pressure signal baseline in the latter part of the exhalation phase. The current composite pressure signal is compared with (i) selected earlier composite pressure signal values and/or (ii) value expected at current time based on extrapolation of composite pressure signal trajectory at specified earlier times and/or (iii) the current rate of change in the composite pressure signal with a selected earlier rates of change. The differences obtained by the comparison are compared with selected threshold values. T onset  is identified when at least one of the differences exceeds the threshold values.

REFERENCE TO RELATED APPLICATION

This application claims priority under 35 USC 119(e) from U.S.Provisional Patent Application No. 60/391,594 filed Jun. 27, 2002.

FIELD OF INVENTION

This invention relates to assisted mechanical ventilation.

BACKGROUND TO THE INVENTION

With assisted ventilation (e.g. assist volume cycled ventilation,pressure support ventilation and proportional assist ventilation)ventilator cycles are triggered by the patient and are intended tocoincide with patient's inspiratory effort. In practice, however, theventilator cycle never begins at the onset of patient's inspiratoryeffort (trigger delay) and the end of the ventilator's inflation phaseonly rarely coincides with the end of inspiratory effort (cycling-offerrors). FIG. 1 provides an example. The bottom channel istransdiaphragmatic pressure (measured by esophageal and gastriccatheters) and reflects true patient inspiratory effort. As may be seen,ventilator cycle was triggered several hundred milliseconds after onsetof effort (interval between vertical lines) and the inflation cyclecontinued well beyond the effort. In fact, the ventilator was cyclingalmost completely out-of-phase with the patient. Trigger delay is oftenso marked that some efforts completely fail to trigger the ventilator(ineffective efforts, e.g. third effort, FIG. 1). A more advanced formof non-synchrony is shown in FIG. 2. In this case, the inflation cycleof the ventilator extends over two patient cycles. There are,accordingly, two inspiratory efforts within a single inflation phase andthere is an additional ineffective effort during the ventilator'sexpiratory phase. The arrows in FIG. 2 indicate the location of theextra patient efforts that did not trigger corresponding ventilatorcycles.

Non-synchrony between patient and ventilator is extremely common. Leunget al found that, on average, 28% of patient's efforts are ineffective(Leung P, Jubran A, Tobin M J (1997). Comparison of assisted ventilatormodes on triggering, patient effort, and dyspnea. Am J Respir Crit CareMed 155:1940-1948). Considering that ineffective efforts are the extrememanifestation of non-synchrony, less severe, yet substantial (e.g. firsttwo breaths, FIG. 1), delays must occur even more frequently.Non-synchrony is believed to cause distress, leading to excessivesedation and sleep disruption, as well as errors in clinical assessmentof patients since the respiratory rate of the ventilator can be quitedifferent from that of the patient. Monitoring respiratory rate is afundamental tool for monitoring critically ill patients on ventilators.

In current mechanical ventilators, triggering occurs when flow becomesinspiratory (i.e. >0) and exceeds a specified amount, or when airwaypressure decreases below the set PEEP (positive end-expiratory pressure)level by a specified amount. Trigger delay has two components. Onecomponent is related to ventilator trigger response and sensitivity.Thus, if the response of the ventilator is poor, triggering may notoccur immediately when the triggering criteria are reached.Alternatively, the threshold for triggering may be set too high by theuser. The component of trigger delay attributable to ventilator responseand sensitivity is given by the interval between zero flow crossing(arrow, FIG. 1) and triggering (second vertical line). The response ofmodern ventilators has improved substantially over the past severalyears such that it is difficult to effect further improvements in thisrespect, and this invention does not contemplate any such improvements.This component of trigger delay can, however, still be excessive if theuser sets an unnecessarily high threshold. This setting may be becauseof lack of sufficient expertise, or because there was excessive baselinenoise at some point, which necessitated a high threshold to avoidauto-triggering. The threshold then remains high even afterdisappearance of the noise.

The second component of trigger delay is the time required, beyond theonset of inspiratory effort (T_(onset)), for expiratory flow to bereduced to zero (interval between first vertical line and the arrow,FIG. 1). This delay is related to the fact that expiratory resistance isusually high in ventilated patients and expiratory time is frequentlytoo short to allow lung volume to return to FRC (functional residualcapacity) before the next effort begins. At T_(onset), therefore,elastic recoil pressure is not zero (DH, dynamic hyperinflation).Inspiratory effort must first increase enough to offset the elasticrecoil pressure associated with DH before flow can become inspiratory,and/or before P_(aw) (airway pressure) decreases below PEEP, in order totrigger the ventilator. By identifying the true T_(onset), which is oneaspect of the current invention, this component of trigger delay(usually the largest component, seen, for example, FIG. 1) can beessentially eliminated.

Cycling-off errors result from the fact that, except with ProportionalAssist Ventilation, current ventilator modes do not include anyprovision that links the end of ventilator cycle to end of theinspiratory effort of the patient. In the most common form of assistedventilation, Volume Cycled Ventilation, the user sets the duration ofthe inflation cycle without knowledge of the duration of patient'sinspiratory effort. Thus, any agreement between the ends of ventilatorand patient inspiratory phases is coincidental. With the second mostcommon form, Pressure Support Ventilation, the inflation phase ends wheninspiratory flow decreases below a specified value. Although the time atwhich this threshold is reached is, to some extent, related to patienteffort, it is to the largest extent related to the values of passiveresistance and elastance of the patient. In patients in whom the product[resistance/elastance], otherwise known as respiratory time constant, ishigh, the ventilator cycle may extend well beyond patient effort, whilein those with a low time constant the cycle may end before the end ofpatient's effort (Younes M (1993) Patient-ventilator interaction withpressure-assisted modalities of ventilatory support. Seminars inRespiratory Medicine 14:299-322; Yamada Y, Du H L (2000) Analysis of themechanisms of expiratory asynchrony in pressure support ventilation: amathematical approach. J Appl Physiol 88:2143-2150).

In U.S. Pat. No. 6,305,374 B1, an approach is described to identify theonset and end of patient's inspiratory effort during non-invasivebi-level positive pressure ventilation (BiPAP). This approach reliesexclusively on the pattern of flow waveform to make theseidentifications. Thus, current values of flow are compared with anestimated value based on projections from preceding flow pattern. If thedifference exceeds a preset amount, a phase switch is declared. Whilethis method may yield reasonably accurate results in the intendedapplication (treatment of obstructive sleep apnea patients withnon-invasive BiPAP), a number of considerations suggest that its use incritically ill, intubated, ventilated patients may not provide accurateresults:

1) Implicit to the use of flow as a marker of respiratory musclepressure output is the assumption that flow pattern reflects changes inalveolar pressure inside patient's lung. This is where respiratorymuscle pressure is exerted. This assumption, however, is true only ifairway pressure is constant. Since airway pressure is one of the twopressure values that determine flow (flow=(airway pressure−alveolarpressure)/resistance), it is clear that changes in airway pressure canalter flow even if there is no change in respiratory muscle pressure. Innon-invasive bi-level support, airway pressure, one of the two pressurevalues that determine flow, is reasonably constant during bothinspiration and expiration, even though the absolute level is differentin the two phases. If one of the two pressure values is constant duringa given phase, it is reasonable to assume that changes in flow duringthat phase reflect changes in the other pressure, namely alveolarpressure. This condition does not apply in intubated, mechanicallyventilated patients. In most modern intensive care ventilators, airwaypressure is actively controlled during expiration through adjustments ofthe PEEP/exhalation valve mechanism. The pattern of such active changesin airway pressure during expiration varies from one ventilator brand toanother and in the same ventilator from time to time depending on thestate of the PEEP/exhalation valve mechanism. Under these conditions,changes in flow trajectory during expiration cannot be assumed toreflect changes in alveolar pressure trajectory. Likewise, duringinspiration airway pressure is far from being constant, regardless ofthe mode used. Thus, changes in inspiratory flow profile cannot be usedto reflect similar changes in alveolar pressure. The use of flow toinfer end of effort during the inflation phase is accordingly notplausible.

2) When passive elastance (E) and resistance (R) are constant over theentire tidal volume range, the product R/E, or respiratory timeconstant, is also constant over the entire period of expiration. Becausethe time constant governs the pattern of lung emptying, a constant R/Eproduces a predictable exponential flow pattern in the passive system.With a predictable pattern it is possible to make forwardextrapolations, or predictions, for the sake of identifying a deviationfrom the expected passive behaviour. Such deviation may then be used,with reasonable confidence, to infer the development of an additionalactive force, such as the onset of inspiratory muscle effort. When E andR are not constant throughout the breath, R/E may change from time totime causing changes in flow trajectory (Δflow/Δt) that are not relatedto muscle pressure. Under these conditions, deviation in Δflow/Δt fromprevious values cannot reliably signify a change in pressure generatedby respiratory muscles. Patients with obstructive sleep apnea, theintended population of U.S. Pat. No. 6,305,374 B1, have generally normallungs; R and E are expected to be constant over the tidal volume range,particularly when expiratory airway pressure is higher than atmospheric(i.e. the usual case when BiPAP is applied). In critically ill,intubated ventilated patients, this is not the case. Resistance is notconstant, primarily because these patients are intubated and theresistance of the endotracheal tube is flow-dependent (the higher theflow, the higher the resistance). The relation between resistance andflow varies from one tube to the other. Furthermore, tidal volume inthese patients often extends into the volume range where elastance isnot constant. Thus, as the lung is emptying, either or both elastanceand resistance may be changing, causing changes in respiratory timeconstant during the same expiration. Under these conditions, changes inflow trajectory need not reflect changes in respiratory muscle pressure.This considerably decreases the sensitivity and specificity of flowpattern as a marker of inspiratory effort.

3) Changes in respiratory muscle pressure (P_(mus)) are not exclusivelyused to change flow. According to the equation of motion, specificallyapplied to intubated patients:P _(mus)=Volume*E+Flow*K ₁+(Flow*absolute flow*K ₂)−P _(aw)  Equation 1

Where, E is passive respiratory system elastance, K₁ is the laminarcomponent of passive respiratory system resistance, K₂ is the resistancecomponent related to turbulence (mostly in the endotracheal tube), andP_(aw) is airway pressure which is determined by the pressure at theexhalation/PEEP valve (P_(valve)), flow and R_(ex), that is resistanceof the exhalation tubing (P_(aw)=P_(valve)−flow*R_(ex)). In thisequation expiratory flow is negative. When P_(mus) changes, as atT_(onset), the flow trajectory should change. However, a change in flowtrajectory also results in changes in volume and P_(aw) trajectories.According to Equation 1, these changes will oppose the change in flow.For example, if expiratory flow decreases at a faster rate, volumedecreases at a slower rate than in the absence of P_(mus). At anyinstant after T_(onset), elastic recoil pressure, which is related tovolume, is higher, and this promotes a greater expiratory flow. The samecan be said for the effect of changes in flow trajectory on P_(aw)trajectory; a lower expiratory flow decreases P_(aw), which promotesmore expiratory flow. How much of the change in P_(mus) is used tochange the flow trajectory depends on the magnitude of the opposingforces. In particular, a higher passive elastance and/or a higher R_(ex)tends to reduce the fraction of the change in P_(mus) used to changeflow trajectory. Furthermore, for a given P_(mus) expended to change theflow trajectory, the actual change in trajectory is determined byresistance (i.e. K₁ and K₂). When E, R_(ex), K₁ and K₂ are all low, amodest change in dP_(mus)/dt results in a sharp change in flowtrajectory. As these characteristics become more abnormal, the change inflow trajectory, for a given dP_(mus)/dt, progressively is attenuated.FIG. 3 illustrates this in a computer simulation.

In the example of FIG. 3, respiratory muscles were inactive in the firstsecond of expiration (as they usually are). This is represented byP_(mus) of zero (lower panel). At 1.0 sec an inspiratory effort begins.P_(mus) rises at a rate of 10 cmH₂O/sec, representative of a normalrespiratory drive. The three flow waveforms represent, from belowupwards, progressively increasing values of K₁, K₂, E and R_(ex). Thevalues used in the lowest waveform are those of a patient with normalpassive elastance and resistance, intubated with a large endotrachealtube (#9 tube, K₂=3), and exhalation tubing with a low resistance(R_(ex)=2). The onset of effort results in a sharp change in the flowtrajectory that can be readily detected within a very short time afterT_(onset).

The middle waveform (FIG. 3) was generated with values representing theaverage intensive care patient on mechanical ventilation. Both passiveK₁ and passive E are higher than normal, K₂ is that of a #8 endotrachealtube, the most common size used, and the exhalation tubing has amoderate (average) resistance. Note that the change in flow trajectoryis considerably less pronounced. An experienced eye, with the benefit ofhindsight (i.e. observing the flow waveform for a substantial periodafter P_(mus) started), may be able to tell that a change in trajectoryoccurred at 1.0 sec. However, it is not possible to prospectivelyidentify that a trajectory change took place in a timely manner, for thesake of triggering the ventilator. Prospective identification of atrajectory change requires comparison between current and previousΔflow/Δt values, or between current flow values and values expectedbased on forward extrapolation of the preceding flow pattern (e.g.dashed lines, FIG. 3). There is always uncertainty with extrapolation,particularly with non-linear functions where the exact function is notknown and, even more so, when the signal is noisy, as the flow signalcommonly is (due to cardiac artefacts or secretions). Comparison ofcurrent and previous Δflow/Δt is also fraught with uncertainties whenthe rate may change for reasons other than respiratory muscle action(see #1 and #2, above). Thus, a wide difference (trigger threshold) mustbe specified, between current and projected flow, or between current andprevious Δflow/Δt, before a trajectory change can be identified withconfidence. Otherwise, false triggering will occur frequently. When thechange in flow trajectory is small, a longer interval must elapse beforethe threshold separation is achieved. It can be seen from the middleflow waveform that a conservative flow separation (between actual andprojected flow) of 0.2 l/sec would not be reached until after flowbecame inspiratory. Thus, in the average mechanically ventilated patientthe use of flow trajectory to identify T_(onset) is not likely to resultin a significant improvement over the current approach of waiting forflow to become inspiratory.

With more severe mechanical abnormalities (top waveform, FIG. 3), thechange in flow trajectory is even more subtle. Even an experienced eye,with the benefit of hindsight, cannot distinguish between a truetrajectory change and some flow artefact. Clearly, with a much strongereffort a flow trajectory change may be identifiable before flow becomesinspiratory. However, when patients have vigorous inspiratory efforts,there is no significant trigger delay even with current triggeringtechniques.

In summary, the use of flow to identify respiratory phase transitions isentirely unsuitable for identification of inspiratory to expiratorytransitions during mechanical ventilation in critically ill patients(because of the highly variable P_(aw) during inflation), and has poorsensitivity and specificity for identifying expiratory to inspiratorytransitions in these patients because of the frequent use of activeexhalation valves, the presence of variable time constant duringexpiration and the often marked abnormalities in elastance andresistance.

SUMMARY OF INVENTION

In one aspect, the present invention provides a method for detecting theonset of inspiratory effort (T_(onset)) in a patient on mechanicalventilation, comprising the steps of:

(a) monitoring airway pressure, rate of gas flow, and volume of gas flowof the patient;

(b) applying a gain factor (K_(f)) to the signal representing rate ofgas flow to convert the gas flow signal into a gas flow pressure signal;

(c) applying a gain factor (K_(v)) to the signal representing volume ofgas flow to convert the gas volume signal into a gas volume pressuresignal;

(d) generating a composite pressure signal (signal) comprising the sumof airway pressure signal, gas flow pressure signal, and gas volumepressure signal, with all signals, having a suitably adjusted polarity;

(e) adjusting K_(f) and K_(v) to result in a desired linear trajectoryof composite pressure signal baseline in the latter part of theexhalation phase;

(f) comparing

-   -   (i) the current composite pressure signal values with selected        earlier composite pressure signal values, and/or    -   (ii) the current composite pressure signal values with values        expected at current time based on extrapolation of composite        pressure signal trajectory at specified earlier times, and/or    -   (iii) the current rate of change in the composite pressure        signal with a selected earlier rate of change in the composite        pressure signal;

(g) comparing differences obtained from such comparison(s) made in step(f) with selected threshold values; and

(h) identifying T_(onset) when at least one of the differences exceedsthe threshold values.

The composite pressure signal may contain a fourth component, consistingof the square of the rate of gas flow to which a gain factor (K_(f2)) isapplied to convert the fourth signal to a pressure signal. K_(f2) mayalso be used to adjust the trajectory of the composite pressure signalbaseline in the latter part of the exhalation phase. K_(f2) may beassigned a value corresponding to the K₂ constant of an endotrachealtube in the patient. The values of K_(v), K_(f) and/or K_(f2) may beadjusted to result in a specified slope or pattern of the compositepressure signal during part or all of the expiratory phase.

A default value of K_(f) may be used while the value of K_(v) isadjusted to obtain a desired baseline composite pressure signaltrajectory. Alternatively, a default value of K_(v) is used while thevalue of K_(f) is adjusted to obtain a desired baseline compositepressure signal trajectory.

The K_(f) or K_(v) value used may be a known or estimated value of therespiratory system resistance or elastance, respectively, of thepatient.

The current composite pressure signal value may be compared with thecomposite pressure signal value at the most recent point where thecomposite pressure signal began a new rising phase and T_(onset) isidentified when the calculated difference exceeds a set threshold value.

T_(onset) detection may be precluded in the early part of the exhalationphase.

The amplitude of the composite pressure signal may be monitored throughthe inspiratory phase and the end of inspiratory effort (T_(end)) isidentified from a reduction in signal amplitude or signal slope below aspecified value, which may be a specified fraction of the highest valueobtaining during the inspiratory phase. T_(end) detection may beprecluded in the early part of the inflation phase. The generatedsignals corresponding to T_(onset) may be used to trigger ventilationcycles and/or signals corresponding to T_(end) may be used to cycle offventilation cycles.

In another aspect of the invention, there is provided a method fordetecting the onset of inspiratory effort (T_(onset)) in a patient onmechanical ventilation, comprising the steps of:

(a) monitoring airway pressure and rate of gas flow of the patient,

(b) applying a gain factor (K_(f)) to the signal representing rate ofgas flow to covert the gas flow signal into a gas flow pressure signal,

(c) generating a composite pressure signal comprising the sum of airwaypressure signal and the gas flow pressure signal,

(d) comparing

-   -   (i) the current composite pressure signal values with values        expected based on extrapolation of composite pressure signal        trajectory at specified earlier times, and/or    -   (ii) the current rate of change of composite pressure signal        with a selected earlier rate of change of composite pressure        signal,

(e) comparing differences obtained from such comparison(s) made in step(d) with selected threshold values, and

(f) identifying T_(onset) when at least one of the differences exceedssaid threshold values.

In this aspect of the invention, the composite pressure signal mayincorporate a third component consisting of the square of the rate ofgas flow, to which a gain factor (K_(f2)) is applied to convert thethird signal to a pressure signal. The selected K_(f) may be a known orassumed value of respiratory system resistance.

The generated signal representing T_(onset) may be used to triggerventilation cycles.

The present invention further includes, methods for determining asuitable threshold value for identifying the onset of inspiratory effortfrom the composite pressure signal obtained according to the proceduresdescribed above.

In one such method, suitable for use where the composite pressure signalincludes the sum of the airway pressure signal, gas flow pressure signaland gas volume pressure signal, and, optionally, the fourth component,comprises:

monitoring the composite pressure signal over suitable intervalspreceding onset of inspiratory effort, in a suitable number of elapsedbreaths;

identifying peaks and troughs in the composite pressure signal over theduration of the intervals;

measuring the changes in signal amplitude between successive peaks andtroughs, the amplitudes reflecting the range of amplitudes of noiseincluded in the composite pressure signal; and

determining from the detected range of noise amplitude, a value thatexceeds the prevailing noise value, such value then being usedprospectively to distinguish between true inspiratory efforts and noise.

Another such method, suitable for use where the composite pressuresignal includes the sum of the airway pressure signal, gas flow pressuresignal and gas volume pressure signal, and optionally, the fourthcomponent, or where the composite pressure signal includes the sum ofthe airway pressure signal and the gas flow pressure signal, andoptionally, the third component, comprising:

monitoring the composite pressure signal over suitable intervalpreceding onset of inspiratory effort in a suitable number of elapsedbreaths;

determining slope of the composite pressure signal in successivesubintervals within the intervals;

measuring the range of slope in the subintervals, such range reflectingthe range of slope change in composite pressure signal related to noise;and

determining from the detected range of slope changes, a difference inslope that exceeds the prevailing noise level, the resulting value thenbeing used prospectively to distinguish between changes in compositepressure signal slope due to inspiratory efforts and those due tocomposite pressure signal noise.

An alternative to the latter method comprises:

monitoring the composite pressure signal over suitable intervalspreceding the onset of inspiratory effort, in a suitable number ofelapsed breaths;

comparing signal amplitude at discrete points within such intervals withvalues predicted to occur at such times from the signal pattern inprevious intervals, the difference in signal amplitude reflecting therange of difference related to composite pressure signal noise; and

determining from the detected range of differences, a value that exceedsthe prevailing noise level, such value then being used prospectively toidentify differences between current and predicted values that reflecttrue inspiratory effort.

In another aspect of the present invention, there is provided a methodfor cycling off the inflation phase of a mechanical ventilator, whichcomprises:

measuring the average interval between successive inspiratory efforts ina patient in a suitable number of elapsed breaths (T_(TOT));

identifying onset of inspiratory effort by utilizing any of theprocedures provided in accordance with the present invention orotherwise;

monitoring the time from the onset of inspiratory effort; and

generating a signal that causes the ventilator to cycle off when timeelapsed since onset of inspiratory effort exceeds a specified fractionof T_(TOT).

The time to generate a signal to cycle off the ventilator may becalculated from the trigger time of current ventilation cycle plus aspecified fraction of T_(TOT).

In a further aspect of the present invention, there is provided a methodfor cycling off the inflation phase of a ventilator in pressure supportventilation, comprising:

measuring the interval between successive inspiratory efforts in asuitable number of elapsed breaths (T_(TOT));

measuring inspiratory flow rate at specified times in the elapsedbreaths which triggered ventilator cycles, the specified timescorresponding to a fraction of the T_(TOT), measured from the onset ofinspiratory effort of each breath or from the trigger time of theventilator;

calculating the average of the flow values obtained at such specifiedtimes in the elapsed breaths; and

generating a signal that causes the ventilator to cycle off wheninspiratory flow in the current inflation phase decreases below saidaverage flow value.

The results concerning patient ventilator interaction may be displayedin suitable format, including but not limited to a monitor, digital orelectrical output ports, or printed material. Such results may include,but not limited to, display of the composite pressure signal, T_(onset)and T_(end) markers and displays regarding trigger delay, cycling-offerrors, patient respiratory rate, number and frequency of ineffectiveefforts, and frequency and duration of central apneas, desirableduration of inflation phase, and flow at a specified fraction of T_(TOT)of the patient in the pressure support ventilation mode.

In accordance with another aspect of the present invention, there isprovided an apparatus for detecting the onset of inspiratory effort(T_(onset)) in a patient on mechanical ventilation, comprising:

circuitry for measuring airway pressure, rate of gas flow and volume ofgas flow of the patient;

amplifier to apply a gain factor (K_(f)) to the signal representing rateof gas flow to convert the signal into a gas flow pressure signal;

amplifier to apply a gain factor (K_(v)) to the signal representingvolume of gas flow to convert the signal into a gas volume pressuresignal;

sunning amplifier that generates a composite pressure signal comprisingthe sum of airway pressure signal, the gas flow pressure signal and thegas volume pressure signal, with all signals having suitably adjustedpolarity;

means to permit adjustment of K_(f) and K_(v) to provide a desiredtrajectory of composite pressure signal baseline in the latter part ofthe exhalation phase;

circuitry to direct the composite pressure signal to a T_(onset)identification circuitry during a suitable period in the expiratoryphase, the identification circuitry comprising circuitry to detect achange in trajectory; and

means for generating a signal corresponding to T_(onset) when measuredchange in composite pressure signal trajectory exceeds a specifiedthreshold.

In the device of the invention, an additional signal may be generated tobe summed by the summing amplifier being generated by multiplying theflow signal by the absolute value of the flow signal and applying a gainfactor (K_(f2)) to the resulting square flow signal using an amplifierand K_(f2) is also used to adjust the trajectory of the compositepressure signal baseline in the latter part of the exhalation phase.K_(f2) may be assigned a value corresponding to the K₂ constant of theendotracheal tube in place in the patient.

The K_(f) value may be fixed at a default value while adjustment ofsignal trajectory is made using K_(v) and/or K_(f2). Alternatively,K_(v) is fixed at a default value while adjustment of signal trajectoryis made using K_(f) and/or K_(f2).

In one embodiment of the invention, the summing amplifier input relatedto volume of flow is omitted.

The device provided herein may include circuitry that precludesT_(onset) identification during an adjustable period after the end ofthe inflation phase of the ventilator.

The T_(onset) identification circuitry may comprise circuitry to obtainthe rate of change of composite pressure signal amplitude and to obtainthe difference between the current rate of change and the rate of changeof the composite pressure signal amplitude at a specified earlier timeand to generate a T_(onset) signal when the difference exceeds a setthreshold value.

The T_(onset) identification circuitry may comprise circuitry to measurethe difference between the current composite pressure signal amplitudeand the composite pressure signal amplitude at a specified earlier timeand to generate a T_(onset) signal when the difference exceeds a setthreshold value.

In the device of the invention, K_(v) and/or K_(f) and/or K_(f2) may beadjusted to produce a horizontal or slightly downward sloping compositepressure signal baseline in the latter part of expiration and theT_(onset) identification circuitry may comprise circuitry to measure thedifference between current composite pressure signal amplitude andcomposite pressure signal amplitude at the most recent point where thecomposite pressure signal began rising and to generate a T_(onset)signal when the difference exceeds a set threshold value.

The composite pressure signal may be gated to circuitry to identify endof inspiratory effort (T_(end)), such circuitry comprising:

circuitry to identify the highest amplitude (peak) of the compositepressure signal reached during the current inspiratory effort;

circuitry to detect when amplitude of the composite pressure signaldecreases below a specified value beyond the time at which the peakoccurred; and

circuitry to generate a signal corresponding to T_(end) when theamplitude of the composite pressure signal decreases below the specifiedvalue, which may be a specified fraction of the peak amplitude of thecomposite pressure signal. Circuitry may be provided to precludedetection of T_(end) during a specified period following ventilatortriggering.

Signal corresponding to T_(onset) maybe used to trigger ventilatorcycles and/or signal corresponding to T_(end) may be used to cycle offinflation phases of the composite pressure signal.

The output of the device may be used for closed-loop control ofventilation setting. Functions executed by electrical circuitry may beexecuted in whole or in part by digital techniques.

In a further aspect of the present invention, there is provided a devicefor estimating a desirable duration of the inflation phase of aventilator, comprising:

circuitry to identify inspiratory efforts of the patient, which may be adevice according to the invention or by other suitable circuitry;

means to calculate the time difference between patient inspiratoryefforts (patient T_(TOT)); and

means for displaying a value corresponding to a specified fraction ofpatient T_(TOT), such specified fraction being a user input or a defaultvalue between 0.3 and 0.5.

In, this device, a signal may be generated to cycle off the inflationphase of the ventilator when the desirable duration has lapsed afterventilator triggering.

A signal may be generated to cycle off the inflation phase of theventilator when the desirable duration has elapsed after onset ofinspiratory effort in current breaths or after a point intermediatebetween onset of effort and ventilator triggering.

A user input may be provided for inputting patient T_(TOT) or itsreciprocal, patient respiratory rate, and the input then is used by thedevice, in lieu of device-determined patient T_(TOT), to determinedesirable duration of inflation phase.

In an additional aspect of the invention, there is provided a device fordetermining the desirable inspiratory flow threshold for terminatinginflation cycles in the pressure support ventilation mode, comprising:

circuitry for estimating desirable duration of inflation phase of theventilator, by using the device provided herein or by any other suitablealternative;

means for measuring inspiratory flow in recently elapsed breaths afterthe desirable duration has elapsed from the ventilator trigger time, orfrom the onset of inspiratory effort preceding triggered breaths, orfrom a specified point in between the two points; and

means for displaying the value of said measured flow.

In such device, the value of the measured flow may be communicated tothe cycling mechanism of the ventilator to effect termination of theinflation phase when the measured flow, or a reasonable approximatethereof, is reached during the inflation phase.

The values relating to patient ventilator interaction determined in thedevices provided herein may be calculated and displayed in suitableformat, including but not limited to a monitor, digital or electricaloutput ports. The values may be any of those discussed above.

The present invention, therefore, concerns a novel method and apparatusto, non-invasively, determine the true onset (T_(onset)) and end(T_(end)) of patient's inspiratory efforts. Such method/device can beused simply as a monitor, informing the user of the presence andmagnitude of trigger delays, ineffective efforts and cycling-off errors.The user can then take appropriate action to reduce the non-synchrony.Alternatively, the method/device can be coupled with the cyclingmechanisms of the ventilator, whereby onset and end of ventilator cyclesare automatically linked to onset and end of patient's efforts, therebyinsuring synchrony without intervention by the user.

One aspect of the current invention is to minimize the cycling-offerrors either by directly identifying the end of patient's inspiratoryeffort or by insuring that the ventilator's inflation phase does notextend beyond the physiologic limit of the duration of inspiratoryeffort.

BRIEF DESCRIPTION OF DRAWINGS

FIG. 1 contains traces of airway pressure, flow and diaphragm pressurefor a patient on mechanical ventilation;

FIG. 2 contains further traces of airway pressure, flow and diaphragmpressure for ventilator cycles;

FIG. 3 is a graphical representation of the effect of variation incertain parameters on change in trajectory of flow upon start ofinspiration;

FIG. 4 is a graphical representation of the effect of variation incertain parameters on change in trajectory of composite pressure signalZ upon start of inspiration;

FIG. 5 contains traces of airway pressure, flow and composite pressuresignal Z calculated in accordance with the invention;

FIG. 6 contains traces of airway pressure, flow, composite pressuresignal Z and diaphragm electrical activity, with the signal Z tracingbeing generated from pressure, flow and volume tracings;

FIG. 7 is a schematic representation of the generation of pressure andflow signals;

FIG. 8 is a block diagram of one embodiment of a device operating inaccordance with the method of the invention;

FIG. 9 is a schematic representation of the digital implementation ofoutput functions;

FIG. 10 contains traces of composite pressure signal Z and T_(onset)integrator output;

FIGS. 11 and 12 show the electrical circuitry used in apparatus of FIG.8;

FIGS. 13 to 17 contain flow charts for the different functions performedby the output microprocessor shown in FIG. 9;

FIG. 18 is a block diagram of one embodiment of a fully digital devicefor carrying out the method of the invention; and

FIGS. 19 to 21 contain flow charts for the different functions performedby the fully digital device of FIG. 18.

DETAILED DESCRIPTION OF THE INVENTION

The present invention contemplates novel methods and devices forspecific and timely identification of respiratory phase transitionswithin the patient for use in monitoring patient-ventilator interactionor to effect switching of ventilator cycles. These methods/devicesrepresent a progression in complexity that address the problems inherentin the prior art ventilation procedures described above.

In the simplest of these methods, a signal is generated (signal X) thatincorporates changes in both the flow and airway pressure (P_(aw))signals. Thus,Signal X=(Flow*K _(f))−P _(aw)  Equation 2,

where, K_(f) is a constant that converts flow to pressure. K_(f) may bean estimated or assumed value of patient's resistance (includingendotracheal tube). There are two advantages to this approach: First,the signal becomes relatively immune to changes in flow trajectoryproduced via changes in pressure at the exhalation/PEEP valve mechanism(#1 in Background above). Thus, if pressure at the exhalation/PEEP valveincreased near the end of expiration (to maintain PEEP), flow willdecrease at a faster rate. Without the P_(aw) component, this effect mayappear as an inspiratory effort. With inclusion of P_(aw) in the signal,changes in flow and P_(aw) tend to cancel out. The extent to which thiscompensation is complete depends on how close K_(f) is to actual patientresistance. In the absence of a known value, a default value may beused, for example 15 cmH₂O/l/sec, representing average resistance(including ET tube) in critically ill, mechanically ventilated patients.With such a default value, correction is not perfect, but the signal ismore specific (than flow) in reflecting T_(onset). Second, by includingP_(aw) in the signal, the signal incorporates that component of P_(mus)that was dissipated against R_(ex) (see #3 in Background). For example,if P_(aw) decreases at T_(onset) (because of the lower expiratory flow),this decrease is summed with the component related to flow, resulting ina sharper change in signal trajectory. With this approach, however,signal baseline prior to inspiratory effort is not flat, but, as in thecase of flow, rises in a non-linear fashion. Forward extrapolationcontinues to be required to identify phase transition. Thus, theuncertainty associated with forward extrapolation is not eliminated butthe change in signal trajectory is sharper, resulting in a more timelydetection of T_(onset) for the same selected detection threshold (i.e.difference between actual and predicted signal required foridentification). Furthermore, this approach continues to be unsuitablefor detection of inspiration to expiration transitions (T_(end)).

A further improvement is achieved by incorporating a component relatedto volume in the signal (signal Y). Thus:Signal Y=Volume*K _(v)+Flow*K _(f) −P _(aw)  Equation 3,

where, K_(v) is a factor that converts volume to pressure. With thistreatment, the increase in the flow term during expiration (note thatflow is negative) is offset by the decrease in the volume term. Thistends to linearize, and decrease the slope of (flatten) the signal inthe interval prior to T_(onset), reducing the uncertainty associatedwith extrapolation, while the change in trajectory at T_(onset) isrendered more acute on account of incorporating representation of allactions resulting from the change in P_(mus) (see #3 in Background). Inthe best case scenario, where K_(v) is identical to passive elastance,K_(f) is identical to passive resistance, and there are nonon-linearities in the passive pressure-flow and pressure-volumerelations, signal Y would be identical to the actual P_(mus) waveform,with a flat baseline and a crisp rising phase at T_(onset) (i.e. as inthe P_(mus) panel of FIG. 3). Under these conditions, extrapolation isunnecessary, and phase transition is identified when signal Y exceeds aset threshold above the baseline value, to account for random baselinenoise. Unfortunately, however, precise determination of actual passiveproperties during assisted ventilation is impossible, and there arenon-linearities in the pressure-flow and pressure-volume relations.These result in some instability in, baseline, necessitating the use ofextrapolation. It may be expected, however, that the transition frombaseline to active inspiration will be crisper after including a volumecomponent (see below).

A further improvement is achieved by allowing for non-linearity in thepressure-flow relation. In mechanically ventilated patients, thenon-linear element is almost exclusively due to endotracheal tubecharacteristics. Thus, a suitable alternate approach is to partition theflow component in two parts, one related to the endotracheal tube andthe other related to a laminar component of resistance (K_(f)). Suchsignal is referred to as signal Z. Thus:Signal Z=Volume*K _(v)+Flow*K _(f)+(Flow*absolute flow*K _(f2))−P_(aw)  Equation 4,

where K_(f2) may be the commercially available K₂ value of theendotracheal tube in place. This treatment essentially eliminates anyartifactual baseline instability related to non-linear pressure-flowbehaviour, further reducing the need for extrapolation and enhancing thecrispness of the transition.

As indicated earlier, precise estimates of E and K₁ are impossible toobtain during assisted ventilation. Passive E and R (including K₁) maybe available from earlier determinations in which the patient was madepassive. These values may be different from the current values, eitherbecause the ventilation conditions under which measurements were madewere different, or true E and R (i.e. K₁) may have changed in theinterim. Some techniques can be used to estimate E and R duringconventional assisted ventilation, but these are not very reliable. Animportant issue, therefore, is the impact of differences between theK_(v) and real E, and between K_(f) and real resistance, on the baselineof the generated signals and on the sharpness of the transition.

In FIG. 4, the same P_(mus) waveform shown at the bottom of FIG. 3 wasused to generate flow, volume and P_(aw) waveforms using valuesrepresentative of the average patient (K₁=10, K₂=5.5, E=25, R_(ex)=5,similar to the values used to generate the middle flow panel of FIG. 3).Signal Z was then generated from the resulting flow, volume and P_(aw)waveforms using inaccurate values of K_(v) and K_(f) (i.e. K_(v)different from real E and K_(f) different from true K₁). Simulationswere made with errors in either direction (over- or underestimation) ofa magnitude that reflects reasonable outside limits of such errors inpractice (i.e. E and K₁ overestimated by 100% or underestimated by 70%).

As may be expected, when there are no errors (i.e. K_(v)=E and K_(f)=K1,middle line, FIG. 4), signal Z is identical to the actual P_(mus)waveform. However, when there are differences between assumed values andactual values, the baseline, prior to T_(onset), is neither flat norlinear. When K_(v) is >E, or K_(f) is <K₁ (upper two lines), baseline issloping down. Under these conditions, there is a qualitative change indirection of signal Z at T_(onset) of effort. Such a directional changecan be easily detected (e.g. by differentiating signal Z and looking,for the point at which the differentiated signal becomes positive).However, when K_(v) is <E, or K_(f) is >K₁ (bottom two lines, FIG. 4),baseline is sloping up and T_(onset) is evident as a change in slope; aquantitative, as opposed to the qualitative, difference observed withthe opposite errors. To identify inspiratory effort under theseconditions, as in the case of flow (FIG. 3), requires forward projectionor extrapolation with the attendant increase in uncertainty and thenecessity to increase trigger threshold. It should be noted, however,that with this approach (i.e. using signal Z (or Y) as opposed to flow)the change in trajectory is much sharper than in the case of flow(middle line, FIG. 3), making it possible to identify inspiratory effortsooner. It should also be noted that the upward slope of the signal,once effort begins, is related to the K_(f) value, being higher whenK_(f) is higher than K₁, and vice versa.

It follows that the use of known values of E and K₁, obtained fromprevious direct measurement, offers advantages over the use of flow.However, under some conditions (i.e. baseline sloping upward)extrapolation techniques (or comparisons between current and previousrates of signal change) are required, and this may delay detection ofphase transition.

A further novel aspect of this invention is to completely ignore patientvalues of E and K₁ and to simply select empiric values of K_(v) andK_(f) that result in a flat or slightly downward sloping baseline in thelatter part of expiration. It is clear from FIG. 4 that, with respect tobaseline pattern (i.e. pattern prior to inspiratory effort), errors canbe made to cancel out. Thus, overestimation of E and overestimation ofK₁ produce opposite errors. If empiric values of K_(v) and K_(f), thatmay have no bearing on actual values, are used, the baseline may besloping up or down depending on the nature and magnitude of errors. Eventhough one cannot tell which value is in error, or by how much, it isalways possible to obtain a flat baseline by adjusting either K_(f) orK_(v). For example, if using the empiric values results in an upwardsloping baseline, the baseline can be made flat by increasing theempiric K_(v) or decreasing the empiric K_(f). If such adjustmentsresult in a flat baseline but some systematic non-linearities persist,these can be offset by adjustments of the non-linear K_(f2) term, ifsignal Z is used, resulting in a flat, and linear baseline. Under suchconditions, identification of T_(onset) presents little difficulty. Aparticularly suitable approach for generating signal Z is to use adefault K_(f) value of 10 cmH₂O/l/sec (15 if signal Y is used) andadjust K_(v) to obtain a flat signal baseline. Alternatively, a defaultK_(v) value (e.g. 25 cmH₂O/l, representing average elastance in ICUpatients) is used and K_(f) is adjusted to obtain a flat signalbaseline. The former approach was found preferable by the inventor as itguarantees a fairly brisk rate of signal rise at T_(onset). Adjustmentsof K_(v) at a set K_(f), or vice versa, can be implemented by the useremploying external inputs for K_(v) and/or K_(f), with feedback from agraphic display of the generated signal (signal Y or Z). Alternatively,selection of the optimum K_(v) and K_(f) values may be doneautomatically using appropriate software.

The above approach does not address the possibility of non-linearpassive pressure-volume relation in the tidal volume range (i.e.non-constant elastance). When this is present, and it is common inmechanically ventilated patients, the respiratory system is stiffer inthe higher part of the tidal range. When K_(v), which is a constant, isadjusted to produce a flat or slightly decreasing signal in the latterpart of expiration the signal is not flat in the early part ofexpiration. In the presence of non-constant elastance (higher elastanceat higher volumes) the signal shows a rising phase in the early part ofexpiration that continues until volume reaches the range of constantelastance. This artifactual rising phase may cause false identificationof a new inspiratory effort. This problem is averted by “blinding” theT_(onset) detection circuitry to the signal during the early part ofexpiration. This can be done, for example, by gating the signal to theT_(onset) detection circuitry only after a certain delay from onset ofexpiratory flow (T_(onset) window delay). Alternatively, the T_(onset)detection circuitry may continue to detect T_(onset) during this periodbut the resulting identification is gated out during this period.Detection of these false triggers can be easily recognized visually bytheir consistent relation to end of ventilator cycle. The magnitude ofthe delay (blinding or blanking period) can then be adjustedaccordingly. Alternatively, software algorithms can be developed todetect triggering signals with a consistent relation to end ofventilator cycle and automatically adjusting the width of the window.

The approach of blinding the T_(onset) detection circuitry to the signalover a time zone close to ventilator cycling-off, where flow is changingrapidly, also helps weed out false triggers related to other artifactsthat commonly occur in the signal at this time (see Cycling-offArtifacts, FIG. 5). These are related to acceleration pressure losses,which are difficult to compensate for, or to phase delays betweenpressure and flow signals, which are common in this setting, among otherfactors.

It should be pointed out that the selected values of K_(v) and K_(f) mayhave little to do with actual patient elastance and resistance. Thesevalues are simply used to facilitate detection of phase transitions. Assuch the actual value of the signal does not reliably reflect actualP_(mus), and such signals cannot be used to reliably estimate the workof breathing or quantitative level of pressure output by the patient.

FIG. 6 shows an example of signal Z generated from pressure, flow andvolume tracings. The signal was generated using a default K_(f) of 10,K_(f2) of 5.5 (ET tube #8) and a K_(v) of 30.5 selected because itproduced a flat baseline in the latter part of expiration. Note the flatbaseline of signal Z in the latter part of expiration. In this patient,diaphragmatic electrical activity was also monitored (lowest tracing),and this reflects the activity of the main inspiratory muscle. Note theexcellent agreement between the onset of effort identified from thesignal Z (arrows) and the onset of diaphragm electrical activity. Notealso that T_(onset) (arrows) was identified much earlier than the timeat which the ventilator triggered with a conventional triggeringalgorithm (T_(trigger), top channel).

A number of approaches can be used to identify a change in signaltrajectory indicative of E→I transition (T_(onset)). Some of theseinclude:

-   -   a) Differentiating the signal (Δsignal/Δt) and comparing current        values with values obtained earlier. T_(onset) is identified        when the difference exceeds a specified amount.    -   b) Comparing current values of signal with predicted values        obtained from forward projection of previous signal trajectory.        T_(onset) is identified when the difference exceeds a specified        amount.    -   c) Comparing current values of signal with values obtained        earlier. T_(onset) is identified when the difference exceeds a        specified amount.    -   d) Preferred approach: Differentiating the signal (Δsignal/Δt)        and identifying points where Δsignal/Δt crosses zero in a        positive direction (t₀(+)). The change in signal amplitude,        relative to amplitude at the immediately preceding t₀(+), is        continuously calculated. T_(onset) is identified when the        difference between current value and value at the preceding        t₀(+) exceeds a specified amount (threshold). If the difference        does not reach threshold by the time Δsignal/Δt crosses zero in        a negative direction (t₀(−)), the difference is reset to zero,        until the next t₀(+). This approach has the advantage of        filtering out slow, random undulations in baseline signal        without altering the relation between signal and inspiratory        effort (which would occur if a simple high pass filter were        used). Such slow, random undulations in baseline signal may be        produced, for example, by changes in thoracic blood volume,        imperfect compensation for mechanical non-linearities, or random        changes in respiratory muscle tone unrelated to phase        transitions. The same approach can also be used to estimate the        amplitude of higher frequency baseline noise (e.g. due to        cardiac artifacts or secretions, see below). Such information        can then be used to automatically adjust the threshold for        identifying T_(onset).

Regardless of which approach is used to identify T_(onset) (a-d, above,or other approaches), a threshold must be set for the magnitude ofchange that must be reached for T_(onset) to be declared. Severalmethods can be used to select such threshold. Some of these include:

-   -   i) A fixed threshold is arbitrarily selected. For example, with        approach (d), a signal increase, beyond the latest t₀(+), of 2        cmH₂O may be used under all conditions. Appropriate values may        be chosen for other approaches. Although feasible, when a        universal threshold is used, the value must be sufficiently high        to avoid false auto-triggering under all circumstances. Since        noise level varies from patient to patient, and from time to        time, such a universal threshold would have to be set to a level        that is unnecessarily high under most conditions.    -   ii) Threshold may be individually selected by the user via        external controls. This can be achieved by the user selecting a        value that results in minimal auto-triggering. Alternatively,        with the help of graphical display of the signal, the user may        adjust the threshold above baseline noise level (e.g. horizontal        dashed line, FIG. 5).    -   iii) Software algorithms can be developed to distinguish noise        from efforts and automatically adjust the threshold accordingly.

The preceding account focussed primarily on identification of E→Itransitions. However, once K_(v) and K_(f) are selected to produce anearly flat baseline during expiration, the shape of the signal duringinspiration (but not necessarily its amplitude, see above) provides areasonable approximation of the shape of inspiratory muscle output(P_(mus)) (for example, see FIG. 6). End of inspiratory effort (T_(end))is normally defined as the point at which inspiratory muscle outputrapidly declines from its peak value. To implement this definition, thehighest value of signal Y (or Z) during the inflation phase can beidentified, in real time, using any of a number of standard techniques.T_(end) is identified when the signal decreases below a specified valueor a specified fraction of peak value.

At times, the signal undergoes a transient artifactual reduction soonafter ventilator triggering. An extreme example is shown in FIG. 5(arrow indicating Ventilator trigger Artifact). It is recognized as anartifact, as opposed to natural end of effort (T_(end)), because thesignal resumes rising again. The presence of these artifacts may causefalse identification of T_(end). To avoid this, if false T_(end)identification occurs, the T_(end) identification circuitry is “blinded”to the signal for a set period after T_(trigger) (see T_(end) WindowDelay, FIG. 5) in the same way the T_(onset) identification circuitry is“blinded” to the signal soon after ventilator cycling-off. Distinctionbetween artifactual and true T_(end) can be easily made by theconsistent occurrence at T_(trigger) and the secondary rise in signalthat characterize false T_(end)s. The distinction can be made by theuser with the help of a monitor displaying the signal, or by usingsoftware algorithms. The width of the T_(end) Window delay is adjustedaccordingly.

At times, true T_(end) occurs soon after ventilator triggering. This isbecause inspiratory muscle activity can be inhibited if inspiratory flowis high, and the ventilator frequently delivers excessive flow soonafter triggering. For this reason, the procedure described above forT_(end) identification may, if used to cycle off the ventilator, resultin medically unacceptable inflation times. A back-up procedure is,therefore, required to insure that the duration of inflation phase isphysiologically appropriate. The same procedure can be used to insurethat the inflation phase does not extend beyond physiologically soundlimits. The following is the rationale and method for ensuring that theduration of the inflation phase remains within physiologic limits.

In spontaneously breathing subjects and patients, the duration of theinspiratory phase (T_(I)) ranges between 25% and 50% of respiratorycycle duration (T_(TOT)). In studies by the inventor using proportionalassist ventilation (PAV), with which the duration of the ventilator'sinflation phase mirrors the patient's own T_(I), the ratio of T_(I) toT_(TOT) (T_(I)/T_(TOT) ratio) was also found to be between 0.25 and 0.5.Therefore, one approach to insure that the duration of the inflationphase is within the physiologic range in modes in which end ofventilator cycle is not automatically synchronized with the patient isto constrain the duration of the inflation phase to be between 0.25 and0.5 of the total cycle duration of patient's own efforts (to bedistinguished from duration of ventilator cycles). Accordingly, inanother aspect of this invention, the end of the ventilator cycle isconstrained to occur within this physiological range. Implementation ofthis procedure requires knowledge of the true respiratory rate of thepatient (as opposed to ventilator rate). The true rate of the patient isthe sum of ventilator rate, the number of ineffective efforts occurringduring the ventilator's exhalation phase (arrows 1 to 3, FIG. 2) and thenumber of additional efforts occurring during inflations triggered by anearlier effort (arrows a to c, FIG. 2). The above-described method foridentifying T_(onset) detects ineffective efforts occurring during theventilator's exhalation phase. These can be added to ventilator rate. Itmay also be possible to identify extra efforts occurring during theinflation phase of the ventilator (a,b,c, FIG. 2) from the generated Yor Z signals. A simpler approach, however, that is particularly suitedfor pressure support ventilation, is to identify points in time at whichflow begins rising again during the inflation phase (FIG. 2). Inpressure support, flow typically declines progressively in the latterpart of inflation. The only possible explanation for a secondary rise inflow, that is sustained for a significant duration (e.g. >0.3 second) isthe occurrence of a second effort during inflation (described byGiannouli et al, American Journal of Respiratory and Critical CareMedicine, vol. 159, pages 1716-1725, 1999). Identification of the extraefforts during the inflation or exhalation phase can be made visually bythe user (FIG. 2). Alternatively, it can be done automatically usingsoftware or analog circuitry. There are several possible approaches toautomatically obtain the number of extra efforts that did not result inseparate ventilator cycles. One such approach is to differentiate theflow signal and determine the number of positive and negative zerocrossings of substantial duration (e.g. >0.4 second, to distinguish fromhigh frequency noise and cardiac artefacts). Another approach is to useFourier frequency analysis of the flow signal. There are clearly othermathematical approaches to identify the characteristic flow transitionsassociated with additional efforts. Thus, it is evident that there aremany ways by which true respiratory rate of patient can be determined.

Once the true respiratory rate of patient is known, it becomes possibleto calculate the real duration of respiratory cycles of the patient(T_(TOT)=60/respiratory rate) and determine the range of inflation timesconsistent with a physiologic T_(I)/T_(TOT). For example, if patient'srate is 30/min, T_(TOT) is 2.0 seconds and the physiologic range for theinflation phase is 0.5 to 1.0 second, reflecting a T_(I)/T_(TOT) rangeof 0.25 to 0.50. Thus, according to this aspect of the invention,average T_(TOT) is determined using any of a number of possible methods.The desirable duration of the ventilator's inflation phase is thendetermined by multiplying T_(TOT) by a user selected physiologicT_(I)/T_(TOT) ratio or a suitable default value (e.g. 0.4). In anotherimplementation of this method, a timer is reset at the onset of a newT_(onset) or a new ventilator cycle. The ventilator ignores othercycling-off commands so long as time elapsed since the last T_(onset),or onset of ventilator cycle, is less than a set value (e.g. 0.3 ofT_(TOT)). Similarly, to guard against excessively long ventilatorcycles, the timer may send a cycling-off command once time, since thelast T_(onset), or onset of inflation phase, exceeds a set fraction ofaverage T_(TOT) (e.g. 0.45). The fractions used for minimum and/ormaximum cycling-off time can be fixed within the ventilator oradjustable by the user.

An adaptation of this last aspect of the invention is particularlysuited for pressure support ventilation (PSV). Because there is oftensome breath by breath variability in T_(TOT), setting the end ofventilator cycle to a fixed fraction of average T_(TOT) results in somecycles having higher, and other cycles having lower, T_(I)/T_(TOT)ratios. In this aspect of the invention, only applicable to PSV, ratherthan causing the ventilator to cycle-off at a predetermined time fromthe last T_(onset), the ventilator is cycled off when inspiratory flowreaches a specified amount, with this specified amount selected toprovide, on average, the specified T_(I)/T_(TOT). This aspect of theinvention is implemented as follows: The interval between successiveinspiratory efforts (T_(TOT)) is determined in several elapsedventilator cycles. The level of inspiratory flow at the specifiedT_(I)/T_(TOT) fraction is noted. For example, if the specified (desired)fraction is 0.4, and T_(TOT) is 3.0 seconds, flow is measured at 1.2second after the preceding T_(onset) set which triggered a ventilatorcycle or, optionally, after the trigger time of the relevant ventilatorcycle. The average of several such determinations, in several elapsedbreaths, is used as the cycling-off flow threshold in subsequentbreaths. With this approach, current cycles destined to have longT_(TOT) automatically have longer inflation cycles. This is so becausethere is normally a correlation between the duration of inspiratorymuscle activity and the T_(TOT) of individual breaths. Thus, in breathsdestined to have a long T_(TOT), inspiratory activity tends to lastlonger and this, in PSV, delays the point at which a specifiedcycling-off flow threshold is reached.

The information provided by the present invention can be utilized in anumber of ways: First, the time of T_(onset), generated by the currentinvention, can be used to trigger ventilator cycles by providing anappropriate signal to the ventilator's triggering mechanism. Second, theend of the ventilator inflation phase can be made to coincide with theend of patient effort, as identified by the present invention, throughappropriate connections to the cycling-off mechanism of the ventilation.Third, cycling-off can be made to occur at specified times or, in thecase of pressure support ventilation, at a specified flow rate, afterT_(onset) or after the onset of ventilator cycle. In this application,the user enters a desired T_(I)/T_(TOT) ratio. The appropriate time, orflow, to cycle-off is then determined from the inputted T_(I)/T_(TOT)ratio and the value of average patient T_(TOT), obtained using thepresent invention. Fourth, cycling off may occur at the identifiedT_(end), conditional on this not violating a specified minimumT_(I)/T_(TOT) ratio.

Whether or not it is used to synchronize the ventilator with patienteffort, the information provided by the current invention can bedisplayed to the user to assist him/her in adjusting ventilator settingsto, indirectly, improve patient ventilator interaction. In thisconnection, the information may be printed out on command or bedisplayed on a monitor. The signal itself can be displayed in real timealong with other useful signals such as flow and airway pressure. Inaddition, numerical values concerning patient ventilator interaction canbe displayed. Some recommended values include:

-   -   a) Trigger delay (difference between ventilator trigger time and        T_(onset)).    -   b) Cycling-off error (difference between ventilator cycling-off        time and end of inspiratory effort).    -   c) True respiratory rate of patient (number of inspiratory        efforts per minute).    -   d) Average duration between inspiratory efforts (T_(TOT)).    -   e) Number of ineffective efforts, per minute or as a fraction of        respiratory rate. This is calculated as the difference between        true rate of the patient and ventilator rate.    -   f) Number of central apneas (no inspiratory efforts for a        specified period, for example 10 seconds) per hour, and/or % of        time spent in central apnea.    -   g) Flow at a specified fraction of average T_(TOT) in the        pressure support ventilation mode.

The numerical values may be accompanied by displayed suggestions on howto adjust ventilator settings to reduce the undesirable aspects ofcurrent interaction.

DESCRIPTION OF PREFERRED EMBODIMENT

The procedures of the present invention as described in details abovemay be implemented in a device which may be constructed as afreestanding device to be attached externally to a ventilator, or may beincorporated within the ventilator. In either case, the operation of thedevice requires inputs related to pressure and airflow in the ventilatorcircuit. FIG. 7 shows a design and components suitable for obtainingthese signals. Although it is possible to obtain these signals byattaching a flow meter and pressure port to the common tube connectingventilator to patient 1, it is preferable to monitor flow and pressureseparately in the inspiratory and expiratory lines and to combine thesignals. This is to avoid clogging of the flow meter and to minimize thenumber of tubing connections extending from near the patient's head tothe device. Accordingly, as shown in FIG. 7, a flow meter and pressureport are inserted in the inspiratory line 2 and another set is insertedin the expiratory line 3. Each set is connected to appropriate pressure4 and flow 5 transducers, which generate signals proportional topressure and flow, respectively. The signals from each pressure 4 andflow 5 transducer is conditioned with suitable low pass filters (e.g. 10Hz) and offset and gain circuitry. Suitable calibrations for thepressure and flow signals are 10 cmH₂O/volt and 1.0 l/sec/volt,respectively. The processed inspiratory 6 and expiratory 7 flow signalsare summed using a summing amplifier 8 to produce a composite flowsignal 9 to be used by the device. The inspiratory 10 and expiratory 11pressure signals are connected to a multiplexer 12. A comparator 13receives the common flow signal 9 and provides a signal 14 to themultiplexer 12 indicating the polarity of the flow signal 9. Themultiplexer generates a pressure signal 15 composed of the inspiratorypressure signal 10 when flow is expiratory and the expiratory pressuresignal 11 when flow is inspiratory. In this fashion the pressure 15measured at any instant is a close approximation of pressure in thetubing near the patient 1 since at all times a static air column existsbetween the active transducer and the common ventilator tubing 1 nearthe patient.

Pressure and flow signals are routinely generated in modern ventilatorsusing an approach similar to that of FIG. 7. If the device of thisinvention is incorporated in the ventilator, the pressure and flowsignals generated independently by the ventilator can be used instead.

FIG. 8 is a block diagram of an analog embodiment of the invention. Asumming amplifier 16 combines four signals, namely a) the pressuresignal 15 suitably inverted 17 and b) the flow signal 9 after suitableamplification 18 using a variable gain amplifier 19. This amplifier 19provides the desired value of K_(f) (Equations 2, 3 and 4). c) Asuitably conditioned and amplified volume signal 20 generated byintegrating 21 the flow signal (9) after subtracting a highly filtered(using an ultra low pass filter 22) flow signal 23 to minimize volumedrift. A variable gain amplifier 24 provides the desired amplification(K_(v)) of the volume signal (as per Equations 3 and 4). d) A signal 25comprised of the product [flow*absolute flow] after suitableamplification. This is an optional signal to be included if it isdesired to compensate for non-linearities in the pressure-flow relationas per Equation 4. The signal corresponding to [flow*absolute flow] isgenerated by processing the flow signal 9 through an absolute valuecircuit 26, and multiplying the output of this circuit 27 by the flowsignal 9 using an analog multiplier circuit 28. The resulting signal 29is then amplified with a variable gain amplifier 30 that provides thedesired value of K_(f2) (Equation 4).

The signal 31 generated by the summing amplifier 16 is further processedby two circuits, one for detecting the onset of inspiratory effort(T_(onset) identification circuit 32) and one for detecting the end ofinspiratory effort (T_(end) identification circuit 33). The overallpurpose of the first circuit 32 is to measure the increase in theamplitude of the signal 31 during periods in which the signal 31 isrising, within a specific time window in the breath determined by aT_(onset) window circuit 34. This time window begins after a specifieddelay 35 from the point at which expiratory flow decreases below aspecified value (e.g. −0.2 l/sec) during expiration. As seen in thediagram of the first circuit 32, the signal 31 is differentiated using adifferentiator 36. The differentiated signal 37 is filtered using anappropriate low pass filter (e.g. 5 Hz) 38 to remove high frequencynoise. The filtered differentiated signal 39 is passed through twocomparators. One comparator 40 sends an enabling positive signal 41 whenthe filtered differentiated signal 39 is positive and the othercomparator 42 sends an enabling positive signal 43 when the filtereddifferentiated signal 39 is negative. The unfiltered signal 37 isintegrated 44 when two gates 45,46 are enabled. The first gate 45 isenabled when the filtered differentiated signal 39 is positive, asdetected by the positive comparator 40. The second gate 46 is enabledduring the specified time window during expiration, as detected by theT_(onset) window circuit 34 and conveyed to the gate by an enablingsignal 47. The integrator 44 is reset whenever the filtereddifferentiated signal 39 becomes negative as detected by the negativecomparator 42. In this fashion integration begins anew only when thesignal is rising within the specified time window. The integrator output48 is received by a comparator 49 which sends out a signal 50,indicating T_(onset), when integrator output exceeds a specifiedthreshold set by an external EI threshold adjust 51.

The specific design used for detection of onset of effort in thisimplementation 32 is selected because it offered an optimal combinationof sensitivity and specificity (i.e. sensitive yet not prone to falsetriggering). It is clear, however, that other designs for detecting achange in signal trajectory are possible. For example, the filtereddifferentiated signal 39, representing current rate of change in signal,can be delayed by a specified amount (e.g. 200 msec). A comparator (notshown) compares the current and delayed forms of the filtereddifferentiated signal. A signal, indicating onset of effort, isgenerated when the difference exceeds a threshold value. Alternatively,the signal itself 31 may be delayed by a specified amount (e.g. 200msec). A comparator (not shown) compares the current and delayed formsof the actual signal and generates a signal, indicating onset of effort,when the difference exceeds a threshold value. Other approaches arepossible within the scope of this invention.

For identifying the T_(end) 33, the signal 31 is first differentiated 52and the differentiated signal 53 is reintegrated 54. The integrator isreset at the onset of inspiratory effort (T_(onset)) using the signal 50generated from the T_(onset) identification circuit 32. In this fashion,any baseline offset in the signal 31 is eliminated and the output of theintegrator 55 reflects only the increase in signal 31 amplitude fromT_(onset). Integrator output 55 is connected to a peak detector circuit56, which is also reset by the T_(onset) signal 50. The output of thepeak detector 57 is attenuated 58 with a suitable attenuation factor(e.g. 50%). Optionally, the attenuation factor may be individuallyadjusted by the user through an external input 59. A comparator 60 sendsa signal 62 when current integrator output 55 decreases below theattenuated peak detector output 61. In this fashion the end ofinspiratory effort is detected when the current integrator output 55decreases below a set percent of the peak level reached during thecurrent inspiratory effort.

At times, the signal 31 or 55 transiently decreases at the time ofventilator triggering (Ventilator Trigger Artifact, FIG. 5). Unlesscorrected, or allowed for, this artefact may result in false detectionof T_(end). A circuit is incorporated to reduce or eliminate theoccurrence of false identification of T_(end). The circuit consists of adelay circuit 63 similar to the one used in the T_(onset) identificationcircuit 34. A timer is activated by a T_(trigger) signal 64 receivedfrom a T_(trigger) identification circuit 65. The latter circuitreceives inputs from the pressure 15 and flow 9 signals. The pressuresignal is differentiated 66 and the resulting signal 67 is directed to acomparator 68 with a suitable reference value (e.g. 15 cmH₂O/sec). Theflow signal 9 also is connected to a comparator 69 with a suitablereference value (e.g. 0.3 l/sec). The outputs of the two comparators68,69 are received by an OR gate 70 which sends a T_(trigger) signal 64to the delay circuit 63 when either the differentiated pressure signal67 or the flow signal 9 exceed the set value in the respectivecomparator 68 or 69. The delay circuit 63 in turns sends a signal 71 toan AND gate 72 after a specified delay set either externally via a userinput 73 or internally as a default value (e.g. 0.2 sec). The AND gate71 also receives the T_(end) signal 62 and sends a final T_(end) signal74 only if it occurs after the specified delay from T_(trigger). In thisfashion, T_(end) signals generated by the triggering artifacts arescreened out.

User Inputs:

The number and types of user inputs may vary depending on howcomprehensive the device is and the extent to which user involvement isdesired by the manufacturer. In the most comprehensive analog embodimentshown in FIG. 8, there are seven user inputs:

-   -   1) K_(f) adjust 75: This input determines the gain of the K_(f)        variable gain amplifier 19. A suitable range is 1 to 25        cmH₂O/l/sec. Because the calibration factors of the flow and        pressure signals may be different (for optimal signal to noise        ratio, see above) an attenuation factor is incorporated to make        allowance for the different calibration factors. For example, if        the flow calibration factor is 1.0 l/sec/volt and the pressure        calibration factor is 10.0 cmH₂O/volt, the relation between the        K_(f) adjust input 75 and the gain of the K_(f) variable gain        amplifier 19 should be 10. In this fashion, the output of the        K_(f) variable gain amplifier, which has units of pressure, is        comparable to the pressure signal 17 at 10.0 cmH₂O/volt.    -   2) K_(f2) adjust 76: This input determines the gain of the        K_(f2) variable gain amplifier 30. A suitable range is 1 to 25        cmH₂O/l²/sec² to take account of the various sizes of        endotracheal tubes used in practice. Again, a suitable        attenuation factor between the K_(f2) input 76 and the actual        K_(f2) gain 30 needs to be incorporated to allow for differences        in pressure and flow calibration factors (see #1 immediately        above).    -   3) K_(v) adjust 77: This input determines the gain of the K_(v)        variable gain amplifier 24. A suitable range is 5 to 100        cmH₂O/l. A suitable attenuation factor between the K_(v) input        77 and the actual K_(v) gain 24 needs to be incorporated to        allow for differences in pressure and flow calibration factors        (see #1 immediately above).    -   4) T_(onset) window delay 35: This input determines the desired        delay, from the point at which expiratory flow decreases below a        set value, before the device begins looking for T_(onset). A        suitable range is 0 to 3.0 seconds.    -   5) E I threshold 51: This determines the amount of increase in        signal amplitude, as detected by the integrator 44 of the        T_(onset) circuit 32, above which T_(onset) is identified. A        suitable range is 0.1 to 10.0 cmH₂O.    -   6) Signal attenuation factor 59: This determines how much signal        amplitude must decrease, after T_(onset), before the T_(end) is        identified. A suitable range is 20 to 90%.    -   7) T_(end) Window delay 73: This input determines the period,        from T_(trigger), during which T_(end) signals 62 are screened        out. A suitable range is 0.0 to 0.3 second.

Some inputs may be deleted by using fixed default values within thedevice. For example, the K_(f) adjust input 75 may be deleted and afixed value of 10.0 is used. A fixed T_(onset) delay value of, forexample, 0.3 second may be used, eliminating the T_(onset) window delayinput 35. A suitable default signal attenuation value (e.g. 50%) may beused replacing the corresponding input 59. Likewise, a T_(end) windowdelay of 0.2 second may be used eliminating the T_(end) window delay 73.Clearly, the more fixed the settings are the less reliable theperformance of the device may become. However, this may be acceptableunder some circumstances with the potential benefit of simplifying theoperation of the device. An alternative would be to have the deviceoperate with default settings unless changed by the user.

Other inputs may also become unnecessary if simpler forms of the signal31 are generated. For example, signal component related to thenon-linear flow function 25 may be eliminated according to Equation 3.In this case the K_(f2) adjust input 76 is deleted. Likewise, the signalcomponent related to volume 20 may be eliminated, according to Equation2, with corresponding deletion of the K_(v) adjust input 77. Again, thesimpler the device, the less reliable its performance will become butthis may be acceptable under certain circumstances. In its simplestform, all the user needs to do is to set the E I threshold input 51.

Device Outputs:

Certain internal signals need to be displayed to allow the user toadjust the input settings, while others provide the user with theresults of monitoring. These signals can be displayed on a monitor 78included in a freestanding device. Alternatively, if the device isincorporated inside the ventilator, the monitor of the ventilator can beused for this purpose. A third embodiment involves directing thedevice's outputs to an analog to digital converter and displaying theoutputs on a separate computer.

The following output signals are necessary for adjusting the inputsettings:

a) The main signal itself 31.

b) The output of the integrator 48 in the T_(onset) circuit (32).

The use of these two signals for the sake of input adjustment isdescribed below under OPERATION (below).

Additionally, the signals representing flow 9, pressure 15 and volume 79may be displayed on the monitor for general monitoring purposes.

Signals representing the onset of inspiratory effort 50 (T_(onset)) andend of inspiratory effort 74 (T_(end)) are also displayed on themonitor. In the event these signals are to be used to actively controlthe cycling of the ventilator, they are communicated to the ventilator'scycling mechanism.

Additional information of value in guiding ventilator setting is mostconveniently generated by a small microprocessor. A block diagram of apreferred embodiment (103) is provided in FIG. 9. Here, the flow signal9 is digitized using an analog to digital converter. In addition, thecentral processing unit receives the signals corresponding to T_(onset)50 and T_(end) 74 of inspiratory effort and signals reflectingventilator trigger time (T_(trigger)) 64 and cycling off time (T_(off),derived from the T_(onset) Window circuit 34 see also 96 in FIG. 12).The latter two signals may also be obtained directly from theventilator. The user inputs the ventilator mode 88 and the desiredT_(I)/T_(TOT) ratio 89. From these data, the microprocessor calculatestrigger delay (T_(trigger)−T_(onset)) 80 and cycling off delay(T_(off)−T_(end)) 81. The flow signal is differentiated. Additionalinspiratory efforts during the inflation phase 82 are identified whenthe differentiated flow becomes positive after an earlier negative phase(Identify additional efforts function 82, FIG. 9). A “calculate patientrate function” (83, FIG. 9) calculates respiratory rate of patient 83from the sum of number of T_(onset) transitions during expiration 50 inthe last minute and the number of additional efforts during inflation 82in the last minute. The number of ventilator cycles per minute 84 iscalculated from the number of T_(trigger) signals 64 in the last minute.The number of ineffective efforts 85 is calculated from the differencebetween patient respiratory rate 83 and ventilator rate 84. Thisadditional information is then displayed on the monitor. Additionally,with knowledge of patient respiratory rate 83 the average breathingcycle duration (T_(tot), T_(tot)=60/respiratory rate) of the patient canbe calculated. The microprocessor calculates the desirable duration ofventilator cycle 87 (desirable T_(I)=T_(TOT)*desirable T_(I)/T_(TOT))where T_(I)/T_(TOT) is a default value (e.g. 0.4) or a user input 89.The microprocessor also receives a user input indicating the mode ofventilation 88. In the PSV mode, the microprocessor samples flow at thedesired T_(I) in several elapsed cycles (Flow at desired T_(I) function90), and displays the average value on the monitor. The user can takeadvantage of this information (desired T_(I) or Flow at desired T_(I))to adjust ventilator settings to result in optimal T_(I)/T_(TOT).

In another embodiment of the output processor 103 patient respiratoryrate (or T_(TOT)) is inputted to the processor, replacing the “CalculatePatient Rate” function 83. This input is then used to calculate the“Desirable T_(I)” 87 and “Flow at Desired T_(I)” 90. Patient respiratoryrate may be determined by the user from inspection of chest movements orby observing the flow tracing on the monitor, or automatically usingcomputational methods other than the ones described in the aboveembodiment 103.

Operation

When the device is built inside the ventilator, the pressure 15 and flow9 signals are permanently connected to the device. For freestandingsystems, the first step is to connect the flow meters and pressure portsto the inspiratory 2 and expiratory 3 lines close to the ventilator(FIG. 7). The device is turned on. Tracing of the Signal 31 appears onthe screen (FIG. 10). Subsequent steps depend on what inputs areavailable on the device and user preference. For the most comprehensiveanalog embodiment (FIG. 8) the recommended procedure is as follows:

-   -   1) Enter the K_(f2) value 76. This is the K₂ value of the        endotracheal tube in use. A table is provided that states the K₂        values for the range of endotracheal tube sizes used.    -   2) Set the other inputs to default values as follows: K_(f)        75=10; K_(v) 77=25; T_(onset) window delay=10% of respiratory        cycle duration. For example if respiratory rate is 20/min, set        the delay to 0.3 sec; Signal attenuation factor 59=50%; T_(end)        Window Delay 73=0.2 second.    -   3) If the baseline of the signal 31 is flat in the latter half        of expiration (e.g. 91, FIG. 10B) no further adjustment of K_(v)        is necessary. If it is not flat (e.g. 92, FIG. 10A), adjust        K_(v) setting 77 to make it flat or slightly sloping down (e.g.        91, FIG. 10B).    -   4) If it is difficult to have reasonably linear signal        trajectory using the K_(v) adjust input 77 alone, adjust the        K_(f2) adjust input 76 up or down as necessary to minimize        non-linearities.    -   5) The tracing representing integrator output 48 (FIG. 10B)        shows relatively large broad waves, representing inspiratory        efforts 93, FIG. 10B), and smaller, briefer spikes representing        noise (94, FIG. 10B). Set the E I threshold level (51) to be        just above the smaller spikes in several consecutive breaths        (e.g. 95, FIG. 10B).    -   6) Display the T_(onset) 50 and T_(end) 74 on the screen. If        there are frequent T_(onset) signals triggered by noise,        increase the level of the E I threshold. If there are frequent        T_(onset) signals triggered early in expiration, increase the        T_(onset) Window delay 35. If the T_(end) signal occurs too        early or too late during the declining phase of the signal 31,        adjust the signal attenuation factor 59 accordingly. If there        are frequent false triggers of T_(end) at the time of ventilator        triggering, increase the T_(end) Window Delay 73 to eliminate        false triggering of T_(end).

FIGS. 11 and 12 show details of the electrical circuitry used in thepreferred embodiment (FIG. 8). All circuits are powered by a suitable+/−8 volts power supply. In FIG. 11 the circuitry used to generate themain signal 31 (block 80 in FIG. 8) is displayed. The summing amplifier16 with its 4 inputs (17,18,20,25) is shown in the lower right corner ofthe Figure. Circuitry used to process the four inputs prior to thesumming amplifier stage is outlined in boxes bearing the same numbers asin the corresponding components of the block diagram (FIG. 8). Theindividual electrical components in each circuit are identified bystandard electrical symbols and the values of resistors and capacitorsindicated are those used in a properly functioning prototype.

FIG. 12 shows details of he electrical circuitry used for T_(onset)identification 32 and T_(end) identification 33. As in FIG. 11, theindividual electrical components in each circuit are identified bystandard electrical symbols and the values of resistors and capacitorsindicated are those used in a properly functioning prototype. Thespecific function of each circuit and its connections to other circuitshave been described in detail in relation to the block diagram of FIG.8, and the design of each circuit is standard for the purpose intendedin each case. Some of the component circuits need additionalexplanation, however:

T_(onset) Window circuit 34: In this circuit the flow signal isconnected to a Schmitt trigger circuit (left half of the T_(onset)window circuit 34) characterized by hysteresis. With the indicatedvalues of the different circuit components, the Schmitt circuit sendsout a constant voltage (8 volts) whenever flow decreases below −0.2l/sec 96. The signal 96 remains on until flow rises to >0.2 l/sec. Inthis application, the onset of the signal 96 indicates the beginning ofthe exhalation phase and is also used to mark the end of ventilatorcycle (T_(off)). The output of the Schmitt trigger circuit is connectedto a delay circuit with an externally adjustable delay time 35. Theoutput of the delay circuit 97 is received by an AND gate 98. The ANDgate 98 also receives the output of the Schmitt trigger circuit 96directly and sends a signal when the T_(onset) window is open, asindicated by the output of the Schmitt trigger circuit 96 but only afterthe specified delay 35 has elapsed, as indicated by a positive outputfrom the delay circuit 97. In turn, the output of the AND gate 47(referred to as Q signal in FIG. 12) serves multiple functions thatinclude enabling one of the transmission gates 46 in the T_(onset)identification circuit 32.

T_(trigger) 64 detection circuitry: There are many ways by which thetime at which the ventilator was triggered can be detected. In thisembodiment T_(trigger) was detected when the rate of increase inpressure exceeded 15 cmH₂O/second OR flow increased beyond 0.4 l/second.To this end, a differentiator 66 was used to obtain Δpressure/Δt 67.Next, a comparator 68 produces a positive signal 99 when Δpressure/Δt 67exceeds a set value of 15 cmH₂O/second. In another circuit 69 acomparator generates a positive signal 100 when flow (9) exceeds 0.4l/second. Two diodes 101,102 function as an OR gate so that a positivesignal (T_(trigger), 64) is generated when either the Δpressure/Δt orflow exceed the set respective thresholds.

T_(end) Window circuit 63: This circuit has four components. First, theQ signal 47, representing time window for T_(onset) detection, isinverted using an inverter 104. The positive phase of this inverted Qsignal 105 defines the maximum period during which T_(end) can belocated. The second component is an AND gate 106 which receives theinverted Q signal 105 and the T_(trigger) signal 64 and sends a positivesignal 107 when both its inputs are positive. The positive edge of thissignal 107 activates a flip-flop switch 108, which is the thirdcomponent of the T_(end) Window circuit. The flip-flop switch 108 isreset by the Q signal 47. The fourth component is a delay circuit 110with an adjustable external control 111. The delay circuit 110 receivesthe output of the flip-flop switch 109. After the set delay, the delaycircuit 110 sends out a positive signal 112, which persists until thebeginning of the Q signal 47. The output of the delay circuit 112 is oneof the two inputs to the main AND gate 72 which generates the T_(end)signal 74.

The other components of the T_(end) circuit 33, as shown in FIG. 12,include the differentiator 52 and integrator 54 that calculate thechange in the main signal 31 since the onset of the current effort (55,referred to as S′ in FIG. 12). The integrator is reset by the T_(onset)signal 50. The peak detector that determines the highest level of signalS′ 55 reached during the current effort, is shown 56 and is also resetby the T_(onset) signal 50. The output of the peak detector 57 isattenuated with an externally adjustable attenuator 58. Finally, acomparator 60 receives the output of the attenuated peak signal 61 andthe differentiated integrated signal (55, S′) and sends a T_(end) signal62 when the latter 55 decreases below the former 61. The T_(end) signal62 is gated out only if the T_(end) Window is open as indicated by apositive output of the T_(end) Window delay circuit 112. This gatingfunction is performed by an AND gate 72.

The circuitry used in this preferred embodiment is clearly not the onlyway by which the functions and results contemplated by the currentinvention can be implemented. Other circuit designs can be used toaccomplish the same objectives and these are within the scope of thisinvention.

FIGS. 13 to 17 show flowcharts for the different functions performed bythe output microprocessor (FIG. 9). The power on start-up routine (113,FIG. 13) clears the memory and enables the Interrupt Request (IRQ)Process. The IRQ process (114, FIG. 14) is executed at suitableintervals (e.g. every 5 msec). It collects data from various inputs (seeFIG. 9 for inputs), calculates the time derivative of flow and storescollected and derived data in memory. It also checks for the times atwhich T_(trigger), T_(onset), T_(off), and T_(end) occur and storesthese times in memory. Because all these timing inputs are squarefunctions, detection of the times at which these events occurred isbased on a simple comparison of current value with the immediatelypreceding one. If current value is high and preceding value is low, theevent is deemed to have occurred. For example, if current value ofT_(trigger) input is high while the immediately preceding value was low,T_(trigger) is deemed to have occurred then, and so on. Finally, the IRQprocess writes the waveform data and calculated variables to themonitor. The main program loop (115, FIG. 15) performs the variousfunctions identified in the block diagram (FIG. 9) in sequence each timea T_(trigger) is detected. The flow charts of individual functions areshown in individual diagrams bearing the same numbers. In the triggerdelay function (80, FIG. 15), when the difference between currentT_(trigger) and the last T_(onset) is >1.0 second, trigger delay isignored. Thus, the maximum trigger delay allowed is 1.0 second.Situations in which T_(onset) occurs more than 1.0 second beforeT_(trigger) are usually ventilator cycles triggered by the ventilatorand not by the patient. The Calculate Ventilator T_(TOT) function (117,FIG. 15) calculates the interval between current and previousT_(trigger). The cycling off delay function (81, FIG. 16) calculates thedifference between the end of ventilator cycle (T_(off)) and end ofpatient effort ((T_(end)) in current cycle. In the Identify AdditionalEfforts function (82, FIG. 17) the program looks in the interval betweenT_(trigger) and T_(off) of the previous cycle for points at whichΔflow/Δt crosses from negative to positive and stays positive for 300msec. When this occurs, it identifies an additional effort during theprevious ventilator cycle and adds its time to the circular buffer forsubsequent counting. The choice of 300 msec is quite conservative andmay suitably be reduced to 200 msec or even less. In the CalculatePatient Rate function (83, FIG. 17) the program calculates the number ofT_(onset) transitions and number of Additional Efforts during inflationin the one-minute interval before the current T_(trigger). In this chartPTE refers to efforts occurring during the exhalation phase of theventilator (i.e. T_(onset) transitions) and PTAE refers to AdditionalEfforts occurring during the inflation phase of the ventilator. In theDesirable T_(I) calculate function (87, FIG. 16) the average patientcycle duration (T_(TOT)) is calculated from 60/patient respiratory ratecalculated in the preceding function (83). Desirable T_(I) is thencalculated from average patient T_(TOT) and the desirable T_(I)/T_(TOT)as indicated by the desired T_(I)/T_(TOT) input (89, FIG. 9). In theCalculate Target Flow function (90, FIG. 16) the first decision iswhether the mode is pressure support ventilation (PSV). If so, theprogram reads flow at an appropriate time in the immediately precedingventilator cycle. There are a number of options for the appropriate timeat which to measure flow (see next paragraph). Occasionally, the time atwhich to measure flow may occur after the end of the ventilator cycle,where flow is negative (i.e. expiratory). This is the case when therespiratory time constant of the patient (resistance/elastance) is tooshort. A provision is made whereby if flow at the chosen time isnegative it is assigned a value of zero. With certain variables it ispreferable to provide the user with average results as opposed to, or inaddition to, results of individual cycles, which may be quite variable.For this reason individual results of certain variables are stored incircular buffer (e.g. Trigger delay (80, FIG. 15), Cycling off delay(81, FIG. 16), Ventilator T_(TOT) (117, FIG. 15) and Target flow for endof cycle (90, FIG. 16)). The Calculate Averages function (116, FIG. 16)then calculates the average values in a preset number of elapsedbreaths. In the illustrated embodiment (116, FIG. 16), the number ofcycles averaged is 10. However, other numbers may be chosen depending onmanufacturer or user preference. Two other variables are derived fromthese averaged values. Ventilator rate (84, FIG. 15) is calculated from[60/average ventilator T_(TOT) (117)) and the number of IneffectiveEfforts (85, FIG. 16) is calculated from the difference between AveragePatient Rate (83, FIG. 17) and Average Ventilator Rate (84, FIG. 15).

In the illustrated embodiment for calculating target flow to cycle offpressure support ventilation (90, FIG. 16) the point chosen to measureflow was the preceding T_(trigger) plus an interval corresponding todesirable T_(I), with the latter based on desired T_(I)/T_(TOT) andaverage respiratory cycle of patient efforts (87, FIG. 16). There are,however, several other options for selecting the point in time at whichto measure flow. These include, but are not limited to:

-   -   a. Desirable T_(I) is added to the T_(onset) preceding the        previous T_(trigger) instead of adding it to T_(trigger) itself.    -   b. Desirable T_(I) is added to a point in time between previous        T_(trigger) and the preceding T_(onset).    -   c. Desirable T_(I) is calculated from desired T_(I)/T_(TOT)        fraction of the T_(TOT) of the individual patient cycle that        included the previous T_(trigger). This value is then added to        the previous T_(trigger), the preceding T_(onset) or some        intermediate time.

Each of these options has advantages and disadvantages. In practice, thedifference in net result should be small. However, some manufacturers orusers may prefer one or the other or even a completely different option.

The resulting output of such microprocessor (FIG. 9) are displayed on amonitor. The information provided can be utilized by the user to adjustventilator settings in order to optimize patient-ventilator interaction.Alternatively, or in addition, some of the outputs can be channelled tothe cycling mechanism of the ventilator to effect such optimizationautomatically. Of particular utility is the use of information providedby the Desirable T_(I) function 87 to automatically set the duration ofthe inflation phase of the ventilator. Likewise, the output of theTarget Flow for End of Cycle in the pressure support mode 90 can be usedto automatically determine the flow threshold at which the ventilatorcycles off in this mode. Other examples of use of generated datainclude, but are not limited to, increasing the flow threshold forcycling off pressure support when the Cycling off Delay function 81produces large positive values or when the Calculate Ineffective Effortsfunction 85 indicates a large number or fraction of such efforts. Themagnitude of pressure support (i.e. amount of increase in pressure attriggering) may also be automatically decreased in the presence of longtrigger delays, as unveiled by the Calculate Trigger Delay function 80,long and positive Cycling off Delays (per 81) or excessive ineffectiveefforts (per 85). Microprocessor output can thus be used for closed loopcontrol of amplitude and duration of ventilator assist.

Whereas the preferred embodiment described herein utilized electricalcircuitry to generate the Signal and to determine T_(onset) and T_(end),it is clear that any and all the functions executed by electricalcircuitry for the current application can be readily executed by digitaltechnology. FIG. 18 is a block diagram illustrating one embodiment of afully digital device. The device receives the various inputs either viaan A/D converter or directly to the central processing unit (CPU)depending on whether the primary inputs are in digital or analog form.In its most comprehensive form, these inputs include pressure 15, flow9, K_(f) 75, K_(f2) 76, K_(v) 77, T_(onset) window delay 35, T_(end)window delay 73, Signal attenuation factor 59, EI threshold 51, Mode 88and desired T_(I)/T_(TOT) 89. The microprocessor executes some functionsin real time and others in non real time when a T_(trigger) isidentified. The non real time functions are similar to those describedin detail in relation to the output microprocessor of FIG. 9 and theassociated flow charts of FIGS. 13 to 17. These will not be describedfurther. The real time functions are executed at suitable intervals;every 5 to 10 msec being optimal. The timed IRQ process 118 isillustrated in flow chart form in FIG. 19. After reading and storing thevarious inputs, it calculates volume and flow². The rate of change inpressure is calculated for use in the Calculate T_(trigger) function 119and the rate of change in flow is calculated for use in the IdentifyAdditional Efforts function 82. The main Signal is then calculatedaccording to Equation 4 and Signal is differentiated for use in theT_(onset) and T_(end) identification functions 121,123. T_(trigger) isthen looked for using a T_(trigger) calculate function 119 and, iffound, the T_(trigger) flag is set to TRUE. This initiates the non realtime functions. Subsequently, the T_(onset) Window calculate function120 is used to determine whether this window is open and, if so, theCalculate T_(onset) function 121 is processed to determine whether aT_(onset) occurred. Finally, the Calculate T_(end) Window function 122and the Calculate T_(end) function 123 are processed to identify if aT_(end) occurred. The T_(trigger) calculate function 119, T_(onset)Window calculate function 120, Calculate T_(onset) function 121,Calculate T_(end) Window function 122, and the Calculate T_(end)function 123 are illustrated in flow chart format in FIGS. 19 to 21.These charts are self-explanatory particularly in light of the detaileddescription of the same functions in relation to the block diagram (FIG.8) and circuit diagrams (FIGS. 11 and 12) of the analog implementation.

As in the case of the analog implementation, the digital implementationcan be simplified to different degrees depending on user andmanufacturer preferences. The outputs of the device may also be expandedor reduced to meet user needs.

SUMMARY OF DISCLOSURE

In summary of this disclosure, the present invention provides a methodand apparatus for detecting the onset and the end of inspiratory effortin a patient on mechanical ventilation. Modifications are possiblewithin the scope of the invention.

1. A method for detecting the onset of inspiratory effort (T_(onset)) ina patient on mechanical ventilation, comprising the steps of: (a)monitoring airway pressure, rate of gas flow, and volume of gas flow ofthe patient; (b) applying a gain factor (K_(f)) to the signalrepresenting rate of gas flow to convert the gas flow signal into a gasflow pressure signal; (c) applying a gain factor (K_(v)) to the signalrepresenting volume of gas flow, also to convert the gas volume signalinto a gas volume pressure signal; (d) generating a composite pressuresignal comprising the sum of airway pressure signal, gas flow pressuresignal, and gas volume pressure signal, with all signals having suitablyadjusted polarity; (e) adjusting K_(f) and K_(v) to result in a desiredlinear trajectory of composite pressure signal baseline in the latterpart of the exhalation phase; (f) comparing (i) current compositepressure signal values with selected earlier composite pressure signalvalues, and/or (ii) current composite pressure signal values with valuesexpected at current time based on extrapolation of composite pressuresignal trajectory at specified earlier times, and/or (iii) the currentrate of change in the composite pressure signal with a selected earlierrate of change in the composite pressure signal; (g) comparingdifferences obtained from such comparison(s) made in step (f) withselected threshold values; and (h) identifying T_(onset) when at leastone of said differences exceeds said threshold values.
 2. The method ofclaim 1 wherein the composite pressure signal incorporates a fourthcomponent, consisting of the square of the rate of gas flow, to which again factor (K_(f2)) is applied to covert said fourth signal to apressure signal and where K_(f2) is also used to adjust the trajectoryof composite pressure signal baseline in the latter part of theexhalation phase.
 3. The method of claim 2 wherein K_(f2) is assigned avalue corresponding to the K₂ constant of the endotracheal tube in placein the patient.
 4. The method of any one of claims 1 to 3 wherein K_(v),K_(f) and/or K_(f2) are adjusted to result in a specified slope orpattern of composite pressure signal during part or all of theexpiratory phase.
 5. The method of any one of claims 1 to 4 wherein adefault value of K_(f) is used while the value of K_(v) is adjusted toobtain a desired baseline composite pressure signal trajectory.
 6. Themethod of any one of claims 1 to 4 wherein a default value of K_(v) isused while the value of K_(f) is adjusted to obtain a desired baselinecomposite pressure signal trajectory.
 7. The method of any one of claims1 to 4 wherein the K_(f) value used is a known or estimated value ofrespiratory system resistance of the patient.
 8. The method of any oneof claims 1 to 4 wherein the K_(v) value used is a known or estimatedvalue of respiratory system elastance of the patient.
 9. The method ofany one of claims 1 to 8 wherein current composite pressure signal valueis compared with the composite pressure signal value at the most recentpoint where the composite pressure signal began a new rising phase andT_(onset) is identified when the calculated difference exceeds a setthreshold value.
 10. The method of any one of claims 1 to 9 whereinT_(onset) detection is precluded in the early part of the exhalationphase.
 11. The method of any one of claims 1 to 10 wherein compositepressure signal amplitude is monitored through the inspiratory phase andwherein the end of inspiratory effort (T_(end)) is identified from areduction in composite pressure signal amplitude, or composite pressuresignal slope, below a specified value.
 12. The method of claim 11wherein said specified value is a specified fraction of the highestvalue obtained earlier during said inspiratory phase.
 13. The method ofclaim 11 or 12 wherein T_(end) detection is precluded in the early partof the inflation phase.
 14. The method of any one of claims 11 to 13wherein signals corresponding to T_(end) are used to cycle offventilator cycles.
 15. A method for detecting the onset of inspiratoryeffort (T_(onset)) in a patient on mechanical ventilation, comprisingthe steps of: (a) monitoring airway pressure and rate of gas flow of thepatient; (b) applying a gain factor (K_(f)) to the signal representingrate of gas flow to covert the gas flow signal into a gas flow pressuresignal; (c) generating a composite pressure signal comprising the sum ofairway pressure signal and the gas flow pressure signal, with the twosignals having suitably adjusted polarity; (d) comparing (i) the currentcomposite pressure signal values with values expected based onextrapolation of composite pressure signal trajectory at specifiedearlier times, and/or (ii) the current rate of change of compositepressure signal with a selected earlier rate of change of compositepressure signal; (e) comparing differences obtained from suchcomparison(s) made in step (d) with selected threshold values; and (f)identifying T_(onset) when at least one of said differences exceeds saidthreshold values.
 16. The method of claim 15 wherein composite pressuresignal incorporates a third component, consisting of the square of therate of gas flow, to which a gain factor (K_(f2)) is applied to convertsaid third signal to a pressure signal.
 17. The method of claim 15 or 16wherein selected K_(f) is known or assumed value of respiratory systemresistance.
 18. The method of any one of claims 1 to 17 whereingenerated signals representing T_(onset) are used to trigger ventilatorcycles.
 19. A method for cycling off the inflation phase of a mechanicalventilator comprising: measuring the average interval between patientinspiratory efforts in a patient in a suitable number of elapsed breaths(T_(TOT)) with said average being updated at suitable intervals;identifying onset of current inspiratory effort; monitoring time fromsaid onset of inspiratory effort; and generating a signal that causesthe ventilator to cycle off when time elapsed since onset of inspiratoryeffort exceeds a specified fraction of T_(TOT).
 20. The method of claim19 wherein the time to generate a signal to cycle off the ventilator iscalculated from the trigger time of current ventilator cycle plus aspecified fraction of T_(TOT).
 21. A method for cycling off theinflation phase of a ventilator in pressure support ventilationcomprising: measuring the interval between successive inspiratoryefforts in a suitable number of elapsed breaths (T_(TOT)); measuringinspiratory flow rate at specified times in those elapsed breaths whichtriggered ventilator cycles, said specified times corresponding to aspecified fraction of the T_(TOT), measured from the onset ofinspiratory effort of said each breath or from the trigger time of theventilator; calculating the average of the flow values obtained at saidspecified times in said elapsed breaths; and generating a signal thatcauses the ventilator to cycle off when inspiratory flow in the currentinflation phase decreases below said average flow value.
 22. A methodfor cycling off the inflation phase of a ventilator in which theventilator is made to cycle off at a time corresponding to the later ofa) end of inspiratory effort, as determined by a method of any one ofclaims 11 to 13, or, b) the time determined from a method of any one ofclaims 19 to
 21. 23. The method of any one of claims 1 to 22 whereinresults concerning patient ventilator interaction are displayed, suchresults including displays of at least one of the composite pressuresignal itself, T_(onset) and T_(end) markers, and trigger delay,cycling-off errors, patient respiratory rate, number and frequency ofineffective efforts, and frequency and duration of central apneas,desirable duration of inflation phase, and flow at a specified fractionof T_(TOT) of the patient in the pressure support ventilation mode. 24.A device for detecting the onset of inspiratory effort (T_(onset)) in apatient on mechanical ventilation, comprising: circuitry for measuringairway pressure, rate of gas flow and volume of gas flow of the patient;amplifier to apply a gain factor (K_(f)) to the signal representing rateof gas flow to convert said signal into a gas flow pressure signal;amplifier to apply a gain factor (K_(v)) to the signal representingvolume of gas flow, to convert said signal into a gas volume pressuresignal; summing amplifier that generates a composite pressure signalcomprising the sum of airway pressure signal, the gas flow-pressuresignal, and the gas volume pressure signal, with all signals havingsuitably adjusted polarity; means to permit adjustment of K_(f) andK_(v) to provide a desired trajectory of composite pressure signalbaseline in the latter part of the exhalation phase; circuitry to directsaid composite pressure signal to a T_(onset) identification circuitryduring a suitable period in the expiratory phase, said identificationcircuitry comprising circuitry to detect a change in trajectory; andmeans for generating a signal corresponding to T_(onset) when measuredchange in trajectory of composite pressure signal exceeds a specifiedthreshold.
 25. The device of claim 24 wherein an additional signal isgenerated to be summed by summing amplifier, said additional signalbeing generated by multiplying the flow signal by the absolute value ofthe flow signal and applying a gain factor (K_(f2)) to the resultingsquared flow signal using an amplifier and wherein K_(f2) is also usedto adjust the trajectory of composite pressure signal baseline in thelatter part of the exhalation phase.
 26. The device of claim 25 whereinK_(f2) is assigned a value corresponding to the K₂ constant of theendotracheal tube in place in the patient.
 27. The device of any one ofclaims 24 to 26 wherein K_(f) is fixed at a default value whileadjustment of signal trajectory is made using K_(v) and/or K_(f2). 28.The device of any one of claims 24 to 26 wherein K_(v) is fixed at adefault value while adjustment of signal trajectory is made using K_(f)and/or K_(f2).
 29. The device of any one of claims 24 to 28 wherein thesumming amplifier input related to the volume of gas flow is omitted.30. The device of any one of claims 24 to 29 including circuitry thatprecludes T_(onset) identification during a specified period after theend of ventilator's inflation phase.
 31. The device of any one of claims24 to 30 wherein the T_(onset) identification circuitry comprisescircuitry to obtain the rate of change in amplitude of the compositepressure signal and to obtain the difference between current said rateof change with said rate of change of the composite pressure signal at aspecified earlier time, and to generate a T_(onset) signal when saiddifference exceeds a set threshold value.
 32. The device of any one ofclaims 24 to 31 wherein the T_(onset) identification circuitry comprisescircuitry to measure the difference between current amplitude of thecomposite pressure signal and signal amplitude of the composite pressuresignal at a specified earlier time, and to generate a T_(onset) signalwhen said difference exceeds a set threshold value.
 33. The device ofany one of claims 24 to 28 and 30 to 32 wherein K_(v) and/or K_(f)and/or K_(f2) are adjusted to produce a horizontal or slightly downwardsloping composite pressure signal baseline in the latter part ofexpiration and the T_(onset) identification circuitry comprisescircuitry to measure the difference between current amplitude of thecomposite pressure signal and amplitude of the composite pressure signalat the most recent point where the composite pressure signal beganrising, and to generate a T_(onset) signal when said difference exceedsa set threshold value.
 34. The device of any one of claims 24 to 33wherein the generated composite pressure signal is gated to circuitry toidentify end of inspiratory effort (T_(end)) said circuitry comprising:circuitry to identify the highest amplitude (peak) of the compositepressure signal reached during the current inspiratory effort; circuitryto detect when amplitude of the composite pressure signal decreasesbelow a specified value beyond the time at which said peak occurred; andcircuitry to generate a signal corresponding to T_(end) when saidamplitude of the composite pressure signal decreases below saidspecified value.
 35. The device of claim 34 where said specified valueis a specified fraction of peak amplitude of the composite pressuresignal.
 36. The device of claim 34 or 35 wherein circuitry is providedto preclude detection of T_(end) during a specified period followingventilator triggering.
 37. The device of any one of claims 24 to 36wherein signals corresponding to T_(onset) are used to triggerventilator cycles and/or signals corresponding to T_(end) are used tocycle off inflation phase of ventilator.
 38. A device for estimating adesirable duration of inflation phase of a ventilator, comprising:circuitry to identify inspiratory efforts of a patient; means tocalculate the time difference between patient inspiratory efforts(patient T_(TOT)); means for displaying a value corresponding to aspecified fraction of patient T_(TOT), said specified fraction being auser input or a default value between 0.3 and 0.5.
 39. The device ofclaim 38 where a signal is generated to cycle off the inflation phase ofthe ventilator when said desirable duration has elapsed after ventilatortriggering.
 40. The device of claim 38 wherein a signal is generated tocycle off the inflation phase of the ventilator when said desirableduration has elapsed after onset of inspiratory effort in currentbreaths or after a point intermediate between onset of effort andventilator triggering.
 41. The device of claim 39 or 40 wherein a userinput is provided for inputting patient T_(TOT) or its reciprocal,patient respiratory rate, and said input is used by the device, in lieuof device-determined patient T_(TOT), to determine desirable duration ofinflation phase.
 42. A device for determining the desirable inspiratoryflow threshold for terminating inflation cycles in the pressure supportventilation mode comprising: circuitry for estimating desirable durationof inflation phase of the ventilator; means to measure inspiratory flowin recently elapsed breaths after said desirable duration has elapsedfrom the ventilator's trigger time, or from the onset of inspiratoryeffort preceding triggered breaths, or from a specified point in betweenthe latter two points; and means for displaying the value of saidmeasured flow.
 43. The device of claim 42 wherein the value of saidmeasured flow is communicated to cycling mechanism of the ventilator toeffect termination of the inflation phase when said measured flow, or areasonable approximation thereof, is reached during the inflation phase.44. The device of any one of claims 24 to 43 wherein values relevant topatient ventilator interaction are calculated and displayed, such valuesincluding displays of at least one of the composite pressure signalitself, T_(onset) and T_(end) markers and displays or outputs indicatingtrigger delay, cycling-off errors, patient respiratory rate, number andfrequency of ineffective efforts, frequency and duration of centralapneas, desirable duration of inflation phase, and flow at a specifiedfraction of patient's T_(TOT) in the pressure support ventilation mode.45. The device of any one of claims 24 to 44 wherein the output of thedevice is used for closed-loop control of ventilator settings.
 46. Thedevice of any one of claims 24 to 45 wherein functions executed byelectrical circuitry are executed in whole or in part by digitaltechniques.